The most common modes of diagnostic ultrasound imaging include B- and M-modes (used to image internal, physical structure), Doppler, and color flow (the latter two primarily used to image flow characteristics, such as in blood vessels). In conventional B-mode imaging, ultrasound scanners create images in which the brightness of a pixel is based on the intensity of the echo return. The color flow mode is typically used to detect the velocity of fluid flow toward/away from the transducer, and it essentially utilizes the same technique as is used in the Doppler mode. Whereas the Doppler mode displays velocity versus time for a single selected sample volume, color flow mode displays hundreds of adjacent sample volumes simultaneously, all superimposed on a B-mode image and color-coded to represent each sample volume's velocity.
Measurement of blood flow in the heart and vessels using the Doppler effect is well known. Whereas the amplitude of the reflected waves is employed to produce black and white images of the tissues, the frequency shift of backscattered waves may be used to measure the velocity of the backscatterers from tissue or blood. The backscattered frequency increases when blood flows toward the transducer and decreases when blood flows away from the transducer. Color flow images are produced by superimposing a color image of the velocity of moving material, such as blood, over the black and white anatomical image. The measured velocity of flow at each pixel determines its color.
The present invention is incorporated in an ultrasound imaging system consisting of four main subsystems: a beamformer 2 (see FIG. 1), a processor subsystem 4, a scan converter/display controller 6 and a master controller 8. System control is centered in master controller 8, which accepts operator inputs through an operator interface (not shown) and in turn controls the various subsystems. The master controller also generates the system timing and control signals which are distributed via a system control bus 10 and a scan control bus (not shown).
The main data path begins with the digitized RF inputs to the beamformer from the transducer. Referring to FIG. 2, a conventional ultrasound imaging system includes a transducer array 50 comprised of a plurality of separately driven transducer elements 52, each of which produces a burst of ultrasonic energy when energized by a pulsed waveform produced by a transmitter (not shown). The ultrasonic energy reflected back to transducer array 50 from the object under study is converted to an electrical signal by each receiving transducer element 52 and applied separately to the beamformer 2.
The echo signals produced by each burst of ultrasonic energy reflect from objects located at successive ranges along the ultrasonic beam. The echo signals are sensed separately by each transducer element 52 and the magnitude of the echo signal at a particular point in time represents the amount of reflection occurring at a specific range. Due to the differences in the propagation paths between an ultrasound-scattering sample volume and each transducer element 52, however, these echo signals will not be detected simultaneously and their amplitudes will not be equal. Beamformer 2 amplifies the separate echo signals, imparts the proper time delay to each, and sums them to provide a single echo signal which accurately indicates the total ultrasonic energy reflected from the sample volume. Each beamformer channel 54 receives the analog echo signal from a respective transducer element 52.
To simultaneously sum the electrical signals produced by the echoes impinging on each transducer element 52, time delays are introduced into each separate beamformer channel 54 by a beamformer controller 56. The beam time delays for reception are the same delays as the transmission delays. However, the time delay of each beamformer channel is continuously changing during reception of the echo to provide dynamic focusing of the received beam at the range from which the echo signal emanates. The beamformer channels also have circuitry (not shown) for apodizing and filtering the received pulses.
The signals entering the summer 44 are delayed so that they are summed with delayed signals from each of the other beamformer channels 54. The summed signals indicate the magnitude and phase of the echo signal reflected from a sample volume located along the steered beam. A signal processor or detector 4 converts the received signal to display data.
The beamformer outputs two summed digital baseband receive beams. The baseband data is input to B-mode processor 4A and color flow processor 4B, where it is processed according to the acquisition mode and output as processed acoustic vector (beam) data to the scan converter/display processor 6. The scan converter/display processor 6 accepts the processed acoustic data and outputs the video display signals for the image in a raster scan format to a color monitor 22.
The B-mode processor 4A converts the baseband data from the beamformer into a log-compressed version of the signal envelope. The B function images the time-varying amplitude of the envelope of the signal as a grey scale using an 8-bit output for each pixel. The envelope of a baseband signal is the magnitude of the vector which the baseband data represent.
The frequency of sound waves reflecting from the inside of blood vessels, heart cavities, etc. is shifted in proportion to the velocity of the blood cells: positively shifted for cells moving towards the transducer and negatively for those moving away. The color flow (CF) processor 4B is used to provide a real-time two-dimensional image of blood velocity in the imaging plane. The blood velocity is calculated by measuring the phase shift from firing to firing at a specific range gate. Instead of measuring the Doppler spectrum at one range gate in the image, mean blood velocity from multiple vector positions and multiple range gates along each vector are calculated, and a two-dimensional image is made from this information. More specifically, the color flow processor produces velocity (8 bits), variance (turbulence) (4 bits) and power (8 bits) signals. The operator selects whether the velocity and variance or power are output to the scan converter 6. Ultimately, the output signal is input to a chrominance control lookup table which resides in the video processor 22.
The acoustic line memories 14A and 14B of the scan converter/display controller 6 respectively accept processed digital data from processors 4A and 4B and perform the coordinate transformation of the color flow and B-mode data from polar coordinate (R-.theta.) sector format or Cartesian coordinate linear array to appropriately scaled Cartesian coordinate display pixel data stored in X-Y display memory 18. In B mode, the intensity data is stored in X-Y display memory 18, each address storing three 8-bit intensity pixels. In color flow mode, the data is stored in memory as follows: intensity data (8 bits), velocity or power data (8 bits) and turbulence data (4 bits).
The scan converter 6 converts the acoustic image data from polar coordinate (R-.theta.) sector format or Cartesian coordinate linear array to appropriately scaled Cartesian coordinate display pixel data at the video rate. This scan-converted acoustic data is then output for display on display monitor 12. In the B mode, the monitor images the time-varying amplitude of the envelope of the signal as a grey scale, i.e., the brightness of a pixel is based on the intensity of the echo return. In the color flow mode, if movement is present, e.g., blood flowing in an artery, a Doppler shift in the return signal is produced proportional to the speed of movements. The display images the flow of blood, i.e., the Doppler shift using different colors, e.g., red for flow toward and blue for flow away from the transducer. In power Doppler imaging, the power contained in the returned Doppler signal is displayed.
Successive frames of color flow or B-mode data are stored in cine memory on a first-in, first out basis. Storage can be continuous or as a result of an external trigger event. The cine memory is like a circular image buffer that runs in the background, capturing image data that is displayed in real time to the user. When the user freezes the system, the user has the capability to view image data previously captured in cine memory. The graphics data for producing graphics overlays on the displayed image is generated and stored in the timeline/graphics processor and display memory 20. The video processor 22 multiplexes between the graphics data, image data, and timeline data to generate the final video output in a raster scan format on video monitor 12. Additionally it provides for various greyscale and color maps as well as combining the greyscale and color images.
Conventional ultrasound scanners create two-dimensional images of a "slice" through an area of the anatomy. Two-dimensional ultrasound images are often hard to interpret due to the inability of the observer to visualize the representation of the anatomy being scanned. However, if the ultrasound probe is swept over an area of interest and two-dimensional images are accumulated to form a three-dimensional volume, the anatomy is easier to visualize. The data may be manipulated in a number of ways, including volume or surface rendering. In addition, the data may be resampled and displayed in planes other than the ones in which the data was originally collected. This allows the user to obtain views of the anatomy that may not be possible given the anatomy and the inability to properly position the probe.
The above techniques have been used to display ultrasound data with varying degrees of success. One problem is that lack of resolution (both spatial and contrast), coupled with speckle and noise in the two-dimensional images make it difficult to properly segment the projected image. The lack of resolution in two-dimensional images is due to a number of factors, including an inability to maintain uniform focus of the beam over a large range, a lack of bandwidth and dynamic range, and a high f-number of the system. Another problem has been the limited range of elevational focus of the beam produced by a single-row, fixed single-focus transducer array. The source data slices used in the reconstruction of a three-dimensional image vary in thickness due to the nonuniform elevation beamwidth. Therefore, the reconstructed images successively degrade as projections or resampled images approach an angle perpendicular to the plane of acquisition. Thus there is a need to lower the f-number and increase the bandwidth of the system, to improve both spatial and contrast resolution of the two-dimensional images, and a further need to control the elevational focus of the ultrasound beam over a greater range, to obtain a much thinner slice of more uniform thickness, enabling improved segmentation in three-dimensional imaging.